Image-forming MR methods which utilize the interaction between magnetic fields and nuclear spins in order to form two-dimensional or three-dimensional images are widely used nowadays, notably in the field of medical diagnostics, because for the imaging of soft tissue they are superior to other imaging methods in many respects, do not require ionizing radiation and are usually not invasive.
According to the MR method in general, the body of the patient to be examined is arranged in a strong, uniform magnetic field (B0 field) whose direction at the same time defines an axis (normally the z-axis) of the co-ordinate system on which the measurement is based. The magnetic field produces different energy levels for the individual nuclear spins in dependence on the magnetic field strength which can be excited (spin resonance) by application of an electromagnetic alternating field (RF field, also referred to as B1 field) of defined frequency (so-called Larmor frequency, or MR frequency). From a macroscopic point of view the distribution of the individual nuclear spins produces an overall magnetization which can be deflected out of the state of equilibrium by application of an electromagnetic pulse of appropriate frequency (RF pulse), so that the magnetization performs a precessional motion about the z-axis. The precessional motion describes a surface of a cone whose angle of aperture is referred to as flip angle. The magnitude of the flip angle is dependent on the strength and the duration of the applied electromagnetic pulse. In the case of a so-called 90° pulse, the spins are deflected from the z axis to the transverse plane (flip angle 90°).
After termination of the RF pulse, the magnetization relaxes back to the original state of equilibrium, in which the magnetization in the z direction is built up again with a first time constant T1 (spin lattice or longitudinal relaxation time), and the magnetization in the direction perpendicular to the z direction relaxes with a second time constant T2 (spin-spin or transverse relaxation time). The variation of the magnetization can be detected by means of one or more receiving RF coils which are arranged and oriented within an examination volume of the MR device in such a manner that the variation of the magnetization is measured in the direction perpendicular to the z-axis. The decay of the transverse magnetization is accompanied, after application of, for example, a 90° pulse, by a transition of the nuclear spins (induced by local magnetic field inhomogeneity) from an ordered state with the same phase to a state in which all phase angles are uniformly distributed (dephasing). The dephasing can be compensated by means of a refocusing pulse (for example a 180° pulse). This produces an echo signal (spin echo) in the receiving coils.
In order to realize spatial resolution in the body, linear magnetic field gradients extending along the three main axes are superposed on the uniform magnetic field, leading to a linear spatial dependency of the spin resonance frequency. The signal picked up in the receiving coils then contains components of different frequencies which can be associated with different locations in the body. The MR signal data obtained via the RF coils corresponds to the spatial frequency domain and is called k-space data. The k-space data usually includes multiple lines acquired with different phase encoding. Each line is digitized by collecting a number of samples. A set of k-space data is converted to a MR image by means of Fourier transformation or other appropriate reconstruction algorithms.
A T2-weighted contrast is often required to characterize tissue lesions detected in MR images (for example in myocardial MR imaging), as the tissue, depending of the type of lesion, has a short T2 relaxation time and thus appears dark in the T2-weighted MR images.
T2-weighted MR images are conventionally acquired using spin echo (SE) or turbo spin echo (TSE) imaging sequences. An alternative would principally be a magnetization prepared turbo field echo (TFE) technique in which a magnetization preparation sequence brings the nuclear magnetization into the transverse plane by an excitation RF pulse, refocuses this transverse magnetization by one or several refocusing RF pulses and finally brings the refocused transverse magnetization back to the z-axis by a corresponding tip-up RF pulse. T2-decay during the period of transverse magnetization, i.e. between the initial excitation RF pulse and the final tip-up RF pulse of the T2 preparation sequence, provides the desired T2 weighting, stored in the z-direction by the tip-up RF pulse. Such a T2 preparation in combination with TFE readout can be designated as T2prep-TFE. T2prep-TFE is known in the art for some special applications, like cardiac/coronary MRI, in which spin echo sequences are less favourable.
However, a problem of the known T2 preparation scheme are interfering signal contributions without T2-weighting. These result from an increasing longitudinal magnetization due to T1 relaxation after the T2 preparation sequence. This non-T2-weighted contamination of the acquired MR signals results in a poor T2 contrast of the reconstructed MR images. The paper ‘Motion and flow insensitive adiabatic T2 preparation modula for cardiac MR imaging at 3 Tesla’ by E. R. Jensita et al. in MRM 70(2013)1360-68 mentions a T2-preparation module that leaves the longitudinal magnetisation in a state that is dependent on its T2.